The subject matter described herein relates generally to pixelated radiation detectors, and more particularly to pixelated radiation detectors for nuclear medical (NM) imaging, such as a radiation camera head for Single Photon Emission Computed Tomography (SPECT), Computed Tomography (CT) and/or Positron Emission Tomography (PET).
In direct conversion detectors, such as semiconductor radiation detectors, the interaction of the incident photons with the semiconductor from which the detector is made produces a cloud of charge-carriers including electron-holes pairs. The electrons drift toward the positively biased pixelated anodes and the holes drift toward the negatively biased cathode. The efficiency of the charge collection by the pixelated anodes depends upon the geometrical structure of the detector and the physical properties of the semiconductor bulk from which the detector is made, among other factors. Additionally, factors not related to the physical properties of the semiconductor bulk, such as lifetime and mobility that affect the detector incomplete charge collection, are the charge sharing between adjacent anodes and the surface recombination in the gap between the anodes.
At least one known method to improve the charge collection is to sum the signal of adjacent pixels in which the signals appear simultaneously. However, the summing method does not solve the problem of the electrons surface-recombination in the gaps between the anodes and also does not recover a large fraction of the shared events that have signals which are below the electronic threshold level. Additionally, the fraction of the shared events that the summing process can account for are reconstructed while increasing the noise (energy resolution).